Ultrasonic diagnostic imaging with contrast agents

ABSTRACT

Apparatus and methods are disclosed for the detection and imaging of ultrasonic contrast agents. Ultrasonic apparatus is provided for coherent imaging of ultrasonic contrast agents, and for detecting harmonic contrast agents. The inventive apparatus includes a dual display for simultaneously viewing a real time image which displays the location of the contrast agent and a triggered contrast image. Methods of contrast agent detection and imaging include the measurement of perfusion rate characteristics, multizone contrast imaging, multifrequency contrast imaging, tissue perfusion display, and high PRF contrast image artifact elimination.

This is a divisional application of U.S. patent application Ser. No.08/723,483, filed Sep. 27, 1996 now U.S. Pat. No. 5,833,613. Thisapplication claims the benefit of U.S. Provisional Application No.60/005,009 filed Oct. 10, 1995; Provisional Application No. 60/013,950filed Mar. 22, 1996; and Provisional Application No. 60/018,095 filedMay 22, 1996.

This invention relates to ultrasonic diagnosis and imaging of the bodywith ultrasonic contrast agents and, in particular, to new methods andapparatus for ultrasonically detecting and imaging with contrast agents.

Ultrasonic diagnostic imaging systems are capable of imaging andmeasuring the physiology within the body in a completely noninvasivemanner. Ultrasonic waves are transmitted into the body from the surfaceof the skin and are reflected from tissue and cells within the body. Thereflected echoes are received by an ultrasonic transducer and processedto produce an image or measurement of blood flow. Diagnosis is therebypossible with no intervention into the body of the patient.

Materials known as ultrasonic contrast agents can be introduced into thebody to enhance ultrasonic diagnosis. Contrast agents are substanceswhich strongly interact with ultrasonic waves, returning echoes whichmay be clearly distinguished from those returned by blood and tissue.One class of substances which has been found to be especially useful asan ultrasonic contrast agent is gases, in the form of tiny bubblescalled microbubbles. Microbubbles present a significant acousticimpedance mismatch in the body, and nonlinear behavior in certainacoustic fields which is readily detectable through special ultrasonicprocessing. Gases that have been stabilized in solutions in the form oftiny microbubbles are infused into the body and survive passage throughthe pulmonary system and circulate throughout the vascular system.Microbubble contrast agents are useful for imaging the body's vascularsystem, for instance, as the contrast agent can be injected into thebloodstream and will pass through the veins and arteries of the bodywith the blood supply until filtered from the blood stream in the lungs,kidneys and liver.

One type of microbubble contrast agent currently under investigationcomprises coated microbubbles. The microbubbles of the contrast agentare covered with a thin biodegradable coating or shell. The microbubbleshave diameters between 0.1 μm and 4.0 μm and a specific density about{fraction (1/10)} of the density of water. The coated microbubbles aresuspended in an aqueous solution for infusion into the blood stream.

Coated microbubbles have the advantage of being stable in the body for asignificant period of time, as the shells serve to protect the gases ofthe microbubbles from diffusion into the bloodstream. The size of themicrobubbles is chosen to enable the microbubbles to pass throughcapillary beds in the body.

At moderately high sound pressure amplitudes the acoustic pressure wavescan cause the shells of coated microbubbles to rupture, freeing thebubbles to behave as noncoated microbubbles until they diffuse into thebloodstream. In their noncoated form acoustic energy can inducenonlinear motion of the microbubbles, itself a detectable ultrasonicphenomenon. This acoustically induced destruction and collapse of themicrobubbles produces a high amplitude response and a characteristicallybright pattern in the color Doppler mode. Hence color Doppler is anadvantageous modality for detecting the collapse of contrast agentmicrobubbles.

U.S. Pat. No. 5,456,257, assigned to the same assignee as the presentinvention, describes a technique for detecting microbubbles throughphase insensitive detection of microbubble destruction anddifferentiation of the detected signals on a spatial basis. Phaseinsensitive contrast agent detection advantageously reduces artifactsfrom moving tissue, and also performs well when imaging contrast agentperfused tissue, where the contrast agent is finely distributed andmoving slowly through the fine capillary structure of tissue. It isdesirable to be able to perform contrast agent imaging with equal effectin large, rapidly moving blood pools such as the chambers of the heart.It is also desirable to specifically tailor the operation of theultrasound machine to harmonic characteristics when performing harmoniccontrast imaging.

In accordance with the principles of present invention, new and improvedapparatus and methods for the detection and imaging of ultrasoniccontrast agents are provided. Ultrasonic apparatus is provided forcoherent imaging of ultrasonic contrast agents, which is advantageous inblood pool contrast imaging. In a second embodiment, the apparatus isspecially tailored to be programmed with response characteristicssuitable for harmonic contrast agents. The inventive apparatus alsoincludes a display for simultaneously viewing a real time image whichdisplays anatomical structures for localization of the contrast agentand a triggered contrast image displaying contrast enhanced images.Methods of employing the inventive apparatus with contrast agentsinclude the measurement of perfusion rate characteristics, multizonecontrast imaging, a technique for discerning larger vessels in a bed offine capillary structures, multifrequency contrast imaging, the displayof contrast enhanced tissue, and a technique for the elimination ofartifacts occurring during high PRF contrast image acquisition.

In the drawings:

FIG. 1 illustrates in block diagram form an ultrasonic diagnostic systemof U.S. Pat. No. 5,456,257 which is capable of performing phaseinsensitive contrast agent detection;

FIG. 2 illustrates in block diagram form an ultrasonic diagnostic systemof the present invention which is capable of performing coherentcontrast agent detection;

FIG. 3 illustrates an ultrasonic image display for contrast agentimaging;

FIG. 4 illustrates in block diagram form a second embodiment of thepresent invention which provides performance advantages for harmoniccontrast agent detection;

FIGS. 5 and 6 illustrate passband characteristics used to explain theperformance of the embodiment of FIG. 4;

FIG. 7 illustrates the principle of time separated pulsing when imagingcontrast agents;

FIG. 8 illustrates an FIR filter structure suitable for use in theembodiment of FIG. 4;

FIGS. 9a-9 d illustrate the effects of stenoses on contrast agentperfusion;

FIG. 10a and 10 b illustrate perfusion curves for good and poorperfusion rates;

FIGS. 11a and 11 b illustrate repetitive perfusion curves for good andpoor perfusion rates;

FIG. 12 illustrates a triggering technique for estimating the perfusioncurve of FIG. 13;

FIG. 13 illustrates a plotted re-perfusion curve.

FIGS. 14a-14 c illustrate a multizone contrast agent scanning technique;

FIGS. 15a and 15 b illustrate display mapping characteristics forcontrast agent imaging;

FIG. 16 illustrates the heart in cross section; and

FIGS. 17a-17 c illustrate the removal of artifacts occurring during highPRF contrast imaging.

Referring to FIG. 1, an ultrasonic diagnostic system described in U.S.Pat. No. 5,456,257 is illustrated in block diagram form. This ultrasoundsystem is capable of performing phase insensitive contrast agentdetection as described in that patent. In the illustrated system,coherent echo signals produced by a beamformer 16 are quadraturedemodulated by an I,Q demodulator 18 to produce quadrature I and Qsignal components. The demodulated signal components are amplitudedetected by an envelope detector 20. The detected signals are filteredby a filter 22 to remove noise and other extraneous signal components.Spatially aligned, temporally separated detected echo signals aredifferentiated by a pulse to pulse differentiation subsystem 24, and thedifferential signals are used to form contrast agent enhanced images.

Performing pulse to pulse differentiation of envelope detected echosignals provides advantages in certain procedures. When contrast agentsare being used in a mode where microbubbles in the fine capillarystructures of tissue are destroyed by ultrasonic waves, differentiationof the echo envelope is particularly useful. In this mode of operation afirst ultrasonic pulse destroys the microbubbles in the tissue and thesedestruction events are received and envelope detected. A second pulse istransmitted to the same locations, and the returning echoes, ideally,show an absence of microbubbles at the locations where the microbubbleswere destroyed. The second set of echoes is subtracted from the firstset on a spatial basis, yielding difference signals of substantialmagnitude at the locations where the microbubbles were destroyed, whichare then displayed at corresponding pixel locations on a display. In arealistic setting the second set of echoes may not actually reveal voidswhere microbubbles were destroyed, due to notional effects, diffusionrates, and other bubble activity. However the difference in bubbleactivity from one pulse to the next will provide a highly detectableresponse when differentiated on a pulse to pulse, spatial basis.

This dramatic difference in scattering characteristics of microbubblesfrom one pulse to another may be due to a host of factors: the burstingof microbubble coatings; oscillation and nonlinear microbubble motion;diffusion of a microbubble during the interpulse interval; or bubblerepositioning, for instance. When microbubble destruction is referred toherein it encompasses the effects of phenomena such as these.

A first embodiment of an ultrasonic diagnostic system constructed inaccordance with the principles of the present invention is shown in FIG.2. This embodiment provides coherent detection of ultrasonic contrastagents. An ultrasonic probe 10 includes an array 12 of ultrasonictransducers which transmit and receive ultrasonic energy. Duringtransmission, an ultrasonic beamformer 16 controls the timing ofactuation of the separate elements of the array 12 by activating thetransducer pursers of a transmitter/receiver 14 at appropriate times topulse the transducer elements so that a steered and focused ultrasonicbeam is produced. During reception ultrasonic echoes received by thetransducer elements are received by transmitter/receiver 14 and coupledto separate channels of the beamformer 16, where the signals areappropriately delayed then combined to form a sequence of coherent echosignals over the depth of reception in the body of the patient.

The coherent echo signals are quadrature demodulated by an I,Qdemodulator 18 which produces quadrature I and Q signal components. Thedemodulated signal components are coupled to a B mode processor 37 whichfilters, detects, and maps greyscale echo signals in the usual manner.The greyscale echo signals of the scanlines of an image are coupled to ascan converter 40 for display of a B mode image.

In accordance with the principles of the present invention the I and Qsignal components are alternatively (or in addition) coupled to a pulseto pulse differentiation circuit 24 which differentiates echoes receivedfrom the same sample volume (location) in the body on a temporal basis.The results of this differentiation are coupled to an amplitude detector20 and the differential response signals are coupled to an eventdiscriminator 27. The event discriminator discriminates events ofmicrobubble destruction at the sample volume location from thedifferentiated echo information. One convenient way to perform thisdiscrimination is by comparison of the detected signals to a thresholdfrom threshold generator 26, passing signals above a threshold andrejecting signals below the threshold. The discriminator will detectmicrobubble destruction events and reject low level noise.

Detected events are coupled to the scan converter 40 for production of aspatial image of the microbubble destruction events in the desired imageformat. The destruction event image may be shown separately, or may becombined with the B mode image to show the contrast agent in relation tosurrounding tissue structure. The images are coupled to a videoprocessor 42 which produces video signals for display on an imagedisplay 50.

This coherent contrast agent detection technique is highly sensitive tosmall variations in microbubble activity in the image area, and performswell when imaging the blood pool of a heart chamber, for instance. In alarge blood pool with a large population of moving microbubbles, theprobability of differential microbubble activity from one pulse toanother is extremely high, explaining the high sensitivity of thistechnique for heart chamber imaging. In comparison with the incoherentcontrast agent detection technique, coherent microbubble detection ismore sensitive to tissue motion and more sensitive to individualmicrobubble events in high contrast agent concentrations. It is alsopossible to process received echoes both coherently and incoherently, toform an image which contains information from both processes.

Contrast agent detection in accordance with the present inventionprovides excellent tissue clutter rejection. Microbubble echo signalsare generally not received alone, but are usually accompanied by echosignals of much greater amplitude which are returned from neighboringtissue and structures. These tissue echoes can be several orders ofmagnitude greater than any of the microbubble echo signals, effectivelymasking them. The pulse to pulse differentiation processing can rejectthe tissue signals by effectively canceling them, revealing the contrastagent echoes which can then be more readily discriminated. Thiscancellation is enhanced by high PRF pulses, which further diminishesmotional artifacts from tissue.

It has been found that the microbubbles of a contrast agent exhibitgreater sensitivity to ultrasonic pulses of certain characteristics andlesser sensitivity to pulses of other characteristics. In general, thehigher the amplitude, the lower the frequency, and (to a lesser extent)the longer the burst length, the more sensitive the microbubbles are todestruction. Thus, the time of occurrence of microbubble destruction canbe modulated and controlled. Microbubbles can be imaged in thebloodstream by scanning at a high frequency with low amplitude, (and toa lesser extent) short burst length pulses. When it is desired tostimulate microbubble destruction, higher power pulses of lowerfrequency and longer burst length are transmitted into the bloodstream.The ultrasound system of the present invention is provided with controlpresets for these two pulse transmission characteristics, enabling theclinician to switch from nondestructive imaging pulses to microbubbledestruction pulses when the clinician so desires. The preferred displayof anatomical structures and microbubble activity employs programmedswitching between the destructive and nondestructive pulse modes forcontrast agent and anatomical structure imaging.

A second embodiment of an ultrasonic diagnostic system constructed inaccordance with the principles of the present invention for use withharmonic contrast agents is shown in FIG. 4. In this second embodimentthe array transducer 112 of the probe 110 transmits ultrasonic energyand receives echoes returned in response to this transmission. Theresponse characteristic of the transducer can exhibit two passbands, onearound the central transmit frequency and another about the center ofthe received passband. For imaging harmonic contrast agents, a broadbandtransducer having a passband encompassing both the transmit and receivepassbands is preferred. The transducer may be manufactured and tuned toexhibit a response characteristic as shown in FIG. 5, in which the lowerhump 60 of the response characteristic is centered about the centertransmit frequency f_(t), and the upper hump 62 is centered about thecenter frequency f_(r) of the response passband. The transducer responsecharacteristic of FIG. 6 is preferred, however, as the single dominantcharacteristic 64 allows the probe to be suitable for both harmoniccontrast imaging and imaging without harmonic contrast agents. Thecharacteristic 64 encompasses the central transmit frequency f_(t), andalso the harmonic receive passband bounded between frequencies f_(L) andf_(c), and centered about frequency f_(r). A typical harmonic contrastagent can have a response such that transmission about a centraltransmit frequency of 1.7 MHz will result in harmonic returning echosignals about a frequency of 3.4 MHz. A response characteristic 64 ofapproximately 2 MHz would be suitable for these harmonic frequencies.

In FIG. 4 a central controller 120 provides a control signal f_(tr) to atransmit frequency control circuit 121 to control the center frequencyand time of transmission of the transmitted ultrasonic energy. Thetransmit frequency control circuit pulses the elements of the transducerarray 112 by means of a transmit/receive switch 114. A preferred methodof pulsing the transducer array is in bursts which scan with sufficientpulses to form an image, followed by intervals of no pulse transmission.Such bursts and intervals are shown in FIG. 7, which shows a burstinterval nPRF and a frame interval t_(Fr), the frame interval includingthe burst interval and an interval of no pulse transmission. The latterinterval allows time for new contrast agent coursing through the body toinfuse the vessels and tissue of the image plane between frame bursts.The frame intervals can be on the order of one second, and can be gatedto the heart rate. During each nPRF burst interval, echoes from the samespatial locations can be gathered for Doppler processing. Preferably ahigh PRF rate such as 6 KHz is used. Imaging procedures of this type arethe subject of U.S. Pat. Nos. 5,685,310 and 5,560,364.

Medical diagnostic ultrasonic scanning is limited by regulatoryrequirements in the peak pressure amplitude of a transmitted pulse andthe integral of the energy transmitted. The preferred scanning ofcontrast agents in accordance with the embodiment of FIG. 4 utilizesrelatively high peak pulse power, with the time integral of transmittedenergy lessened by the intervals during which no pulses are transmitted.The ultrasound system is set to operate with a relatively highmechanical index and an SPTA moderated by the gated or interval bursts.

Echoes received by the transducer array 112 are coupled through the T/Rswitch 114 and digitized by analog to digital converters 115. Thesampling frequency f_(s) of the A/D converters 115 is controlled by thecentral controller. The desired sampling rate dictated by samplingtheory is at least twice the highest frequency f_(c) of the receivedpassband and, for the preceding exemplary frequencies, might be on theorder of at least 8 MHz. Sampling rates higher than the minimumrequirement are also desirable.

The echo signal samples from the individual transducer elements aredelayed and summed by a beamformer 116 to form coherent echo signals.The digital coherent echo signals are then filtered by a digital filter118. In this embodiment, the transmit frequency f_(t) is not tied to thereceiver, and hence the receiver is free to receive a band offrequencies which is separate from the transmitted band. The digitalfilter 118 bandpass filters the signals in the passband bounded byfrequencies f_(L) and f_(c) in FIG. 6, and can also shift the frequencyband to a lower or baseband frequency range. The digital filter could bea filter with a 1 MHz passband and a center frequency of 3.4 MHz in theabove example. A preferred digital filter is a series of multipliers70-73 and accumulators 80-83 as shown in FIG. 8. This arrangement iscontrolled by the central controller 120, which provides multiplierweights and decimation control which control the characteristics of thedigital filter. Preferably the arrangement is controlled to operate as afinite impulse response (FIR) filter, and performs both filtering anddecimation. For example, only the first stage output 1 could becontrolled to operate as a four tap FIR filter with a 4:1 decimationrate. Temporally discrete echo samples S are applied to the multiplier70 of the first stage. As the samples S are applied, they are multipliedby weights provided by the central controller 120. Each of theseproducts is stored in the accumulator 80 until four such products havebeen accumulated (added). An output signal is then produced at the firststage output 1. The output signal has been filtered by a four tap FIRfilter since the accumulated total comprises four weighted samples.Since the time of four samples is required to accumulate the outputsignal, a 4:1 decimation rate is achieved. One output signal is producedfor every four input samples. The accumulator is cleared and the processrepeats. It is seen that the higher the decimation rate (the longer theinterval between output signals), the greater can be the effective tapnumber of the filter.

Alternatively, temporally separate samples are delayed by delay elementsτ and applied to the four multipliers 70-73, multiplied, and accumulatedin the accumulators 80-83. After each accumulator has accumulated twoproducts, the four output signals are combined as a single outputsignal. This means that the filter is operating as an eight tap filterwith a 2:1 decimation rate. With no decimation, the arrangement can beoperated as a four tap FIR filter. The filter can also be operated byapplying echo signals to all multipliers simultaneously and selectivelytime sequencing the weighting coefficients. A whole range of filtercharacteristics are possible through programming of the weighting anddecimation rates of the filter, under control of the central controller.

Returning to FIG. 4, filtered echo signals from tissue, generallyfiltered by a passband centered about or demodulated from the transmitfrequency, are coupled to a B mode processor 37 for conventional B modeprocessing. Filtered echo signals of the contrast agent passband arecoupled to a contrast signal detector 128 which eliminates stationarytissue signals by pulse to pulse subtraction of temporally discreteechoes from a given spatial location, amplitude or envelope detects theresulting difference signals, and discriminates for motion signalcomponents on an amplitude basis. Simple two pulse subtraction of theform P₁−P₂ may be employed where P₁ represents the echoes receivedfollowing one pulse and P₂ represents the echoes received followinganother pulse. Three pulse subtraction of the form |P₁−P₂|+|P₂−P₃| maybe employed to accumulate more signals from successive bubbledestruction pulses.

The filtered echo signals from the digital filter 118 are also coupledto a Doppler processor 130 for conventional Doppler processing toproduce velocity and power Doppler signals. The outputs of theseprocessors are coupled to a 3D image rendering processor 132 for therendering of three dimensional images, which are stored in a 3D imagememory 134. Three dimensional rendering may be performed as described inU.S. Pat. No. 5,720,291, and in U.S. Pat. Nos. 5,474,073 and 5,485,842,the latter two patents illustrating three dimensional power Dopplerultrasonic imaging techniques. The signals from the contrast signaldetector 128, the processors 37 and 130, and the three dimensional imagesignals are coupled to a video processor 140 where they may be selectedfor display on an image display 50 as dictated by user selection. Thevideo processor preferably includes persistence processing, wherebymomentary intensity peaks of detected contrast agents can be sustainedin the image. One technique for providing persistence is through frameaveraging, whereby new image frames are combined with previous frameinformation on a spatial basis. The combination can be done by weightingthe contributions of the old and new frame information and the frameinformation can be combined in a recursive manner; that is, old frameinformation is fed back for combining with new frame information. Apreferred persistence technique is the fast attack, slow decay techniquedescribed in U.S. Pat. No. 5,215,094, which can be applied to bothDoppler and contrast agent images.

Several imaging formats have been found to be preferred for contrastimaging. Power motion imaging as described in U.S. Pat. No. 5,718,229 inwhich the intensity of signals resulting from moving tissue isdisplayed, has been found to be highly diagnostic for structures such asthe walls of the heart when perfused with contrast agents. Power Dopplerimaging has been found to yield excellent results for bloodflow. Threedimensional power Doppler imaging of vessels infused with contrast agentprovide excellent visualization of the continuity of bloodflow andstenoses. The combination of B mode or power motion structuralinformation with power Doppler signals in accordance with thesemi-transparent rendering techniques of the aforementioned U.S. Pat.No. 5,720,291 provides superb renderings of both flow and surroundingstructure.

A preferred display format for contrast agent imaging is depicted by thescreen display of FIG. 3. In this display the signals produced by the Bmode processor 37 are used to display a real time image display 160 ofstructure in the body such as a blood vessel 170. This real time imageis used by the clinician to ascertain and locate the area of the body tobe imaged. Preferably the B mode image is created from echoes returningfrom nondestructive ultrasonic imaging pulses. As discussed above,pulses of low amplitude, high frequency, and short burst duration willgenerally not destroy the microbubbles. However, echoes from pulsesdestructive of microbubbles are used by the contrast signal detector 128to produce contrast agent images 160′ on the same or an adjacentmonitor. Preferably the contrast agent images 160′ are triggered to beacquired at a predetermined phase of the heart cycle, using a heart gatetriggering from the phases of the heartbeat waveform. When the heartbeatis at the desired phase of its cycle, a burst of relatively highamplitude, low frequency, long burst duration pulses are transmitted todestroy the microbubbles in the image plane and detect and display thoseevents. A B mode image acquired at or near the same heartbeat phase isdisplayed, with the vessel or organ 170′ filled in with the imagedmicrobubble destruction events. Thus, the display screen of FIG. 3 willshow a B mode image 160 in real time, and a contrast agent image 160′which is updated each heart cycle.

While the foregoing image presentation is especially useful incardiology where the beating heart is constantly in motion, a variationof this presentation is especially useful in radiology where tissuestructure is more stationary. In the variation, a real time B mode image160′ of anatomical structure is shown, with fluid flow 170′ filled inwith color Doppler. This real time color flow Doppler image is thenperiodically filled in with detected contrast agent, sharplyilluminating the bloodflow. The colorflow Doppler display and thecontrast agent display, both of which are filling in the same areas ofthe anatomical display, may be shown in the same, similar, orcontrasting colors and intensities. The periodicity of the overlaidcontrast agent display may be synchronized to the heart cycle with anEKG trigger as described above, or the periodicity may be chosen by theuser and asynchronous to the heart cycle.

A contrast agent procedure which is advantageously performed inaccordance with the present invention is the measurement of the rate ofperfusion of an organ or area of the body. FIG. 9a illustrates thetravel of an intravenous injection of contrast agent to a capillary bed200. The agent travels in the bloodstream as it moves from the injectionsite 208 and traverses the right ventricle 202, the lungs 204, and theleft ventricle 206 before reaching an artery 209. The contrast agentthen begins to infuse the tissue of the capillary bed 200 as blood flowsfrom the artery 200 through the arterioles 210 and into the capillariesof the tissue.

The perfusion rate into the capillary bed can be used to evaluate theviability of bloodflow in that region of the body or to identify thelocation of a stenosis. Ultrasonic pulses are transmitted to destroymicrobubbles in a region 212 across the capillary bed 200, as shown inFIG. 9b. If a stenosis 214 is impeding the flow of blood in the artery209 and hence to the entire capillary bed 200, the rate of reperfusionof microbubbles will be slow across the entire region 212. But if thestenosis 216 is in an artery which feeds only part of the capillary bed200, the rate of perfusion will be slow in only the portion 218 of theregion which is fed through the stenotic artery. This difference in therate of reperfusion is illustrated graphically by the curves of FIGS.10a and 10 b. Each of these curves shows the same blood volume and hencethe same initial microbubble concentration 220 before the microbubblesare destroyed in the capillary bed. At time td ultrasonic pulses destroythe microbubbles as indicated by the vertical spike in each curve. Whenblood is flowing freely into the capillary bed, a rapid rate ofreperfusion of microbubbles occurs as indicated by curve 222 in FIG.10a. The curve 222 rapidly rises back to the stable microbubbleconcentration level 220. But when the bloodflow is impeded, the rise ofthe curve 224 is much more gradual, as indicated in FIG. 10b. Thereperfusion curve can be repeated continually as shown by FIG. 11a and11 b. FIG. 11a shows a repetitive sequence of reperfusion curves 222,each returning to the full perfusion level 220 in a period of timet_(p). In FIG. 11b, each curve 224 of the same duration t_(p) is shortof the full perfusion level 220 by an amount indicated by arrows B—B.

The reperfusion curve may be reproduced as indicated in FIG. 13.Ultrasonic pulses are transmitted at time t_(d) to destroy themicrobubbles in the capillary bed. A short time later pulses aretransmitted again, the echoes received and imaged to this time measurethe degree of microbubble reinfusion, either by destroying reinfusedmicrobubbles and recording the destruction events, or by counting orintegrating pixels in the area which show reinfused microbubbles. Themeasure of the number of microbubbles reinfused to the region is plottedas a point X of the curve 224. Nondestructive pulses can be repetitivelytransmitted and echoes received to plot a sequences of X points on thecurve as shown in FIG. 13.

Another way to measure the X points on the reinfusion curve throughreadily detectable microbubble destruction events is to utilize a cyclicmeasure similar to the repetitive pattern of FIG. 11b. The cyclicmeasure is useful where the flow in the region is strongly pulsatile dueto the heartbeat cycle. FIG. 12 shows a heart cycle waveform 230,indicating the pulsatile action of bloodflow. At the peaks of thewaveform 230, new blood is pumped into regions of the body during thesystolic phase of the heart cycle. Advantage is taken of this reinfusingaction by repetitively measuring the degree of contrast agent reinfusionat a constant point in the heart cycle, but following continuallydiffering phases of microbubble destruction. In FIG. 12 the X points ofreinfusion measurement all occur at the same phase of the heart cycle.The X points are preceded by changing times at which the microbubblesare destroyed, as indicated by arrows 232, 234, and 236, whichsuccessively precess to earlier times in the heart cycle. This meansthat each X_(n) point of FIG. 12 will be a later X_(n) point on thecurve 224 of FIG. 13. Since the purpose of ultrasonic transmission atthe times of arrows 232, 234, and 236 is to destroy the microbubbles, itis not necessary to receive and analyze the returning echoes at thesetimes. Echo reception and analysis is done at the times of the Xs, andthe Xs shown in FIG. 12 can be plotted as the successive Xs in FIG. 13due to the precession of the destruction time phases indicated by thearrows.

For cardiac imaging it may be desirable to trigger the X_(n) times insynchronization with the diastolic phase of the heart cycle when thecoronary arteries are reinfused with blood. Triggered or gatedacquisition is especially significant in cardiac imaging to reducetissue motion artifacts stemming from the beating movement of the heart.

This technique of measuring perfusion by microbubble destruction canalso be used to image the flow in major vessels of a capillary bed. InFIG. 9d, for instance, it is seen that the major vessels 240 reinfuseearlier than the fine capillaries in a microbubble depleted region 212.The major vessels 240 can be revealed by detecting microbubbles in theregion 212 shortly after pulses have destroyed all of the microbubblesin the region, at which time only the major vessels 240 have beensignificantly reinfused with contrast agent.

It has been found that it is at times not possible to destroy allmicrobubbles in the image plane due to several factors. Sincemicrobubbles are destroyed by high energy, focused ultrasound beams tendto destroy more microbubbles near the beam focal point than at otherlocations. Also, when a dense concentration of microbubbles is to bedestroyed, a great deal of the ultrasonic pulse energy is attenuated bythe near field microbubbles, leaving insufficient energy to destroy farfield microbubbles. A technique for overcoming these effects is shown inFIGS. 14a-14 c. In these drawings, the horizontal axis represents depthinto the body, with the skin line SL indicated at the left side of eachdrawing. A typical ultrasonic image may show the skin line at the top ofthe image and the deepest penetration into the body at the bottom of theimage. To bring a maximum level of energy to bear on the microbubbles inthe image plane, focused pulse are transmitted to focus the ultrasonicenergy on the microbubbles which are to be destroyed. When imaging is tobe done to a significant depth in the body, the pulses will not befocused over the full image depth, but will come into focus around aparticular focal point and then diverge at greater depths. This isindicated in FIG. 14a, where a transmitted pulse is focused at a focalpoint F₁ which is in a focal zone Z₁.

Above this first focal zone Z₁ is a line 270, which represents completemicrobubble destruction over this near field part of the focal zoneZ_(i) and about the focal point F₁. Beyond the focal point the degree ofmicrobubble destruction decreases, as indicated by the declining line272. These lines are shown as straight lines for ease of explanation; itwill be understood that the effect will usually be continually changingand that actual effects may follow a curved relationship.

FIG. 14a represents the transmission of a first pulse along a given beamdirection, a result of which is that near field microbubbles aredestroyed as indicated by lines 270 and 272. Following this microbubbledestruction, a second pulse is transmitted to gather echoes from alongthe microbubble depleted beam direction. The echoes from the two pulsesmay be differentiated and displayed using the ultrasonic apparatus ofFIGS. 1, 2, or 4.

The next pulse transmission for microbubble destruction is focused at asecond focal point F₂ in a second focal zone Z₂ of the beam. Thetransmitted pulse energy will readily reach the second focal zone, sincethe microbubbles in the nearer first zone were previously destroyed.FIG. 14b illustrates this transmission to the second focal zone. Line282 indicates that the remaining microbubbles at the end of the firstzone and the beginning of the second will be destroyed by the seconddestruction pulse, as will microbubbles around the focal point asindicated by line 280. Beyond the second focal point F₂ the degree ofmicrobubble destruction will decline as pulse energy declines, asindicated by line 284. A second interrrogation pulse may be transmittedfollowing the second destructive pulse to differentially detect thesecond sequence of microbubble destruction events.

Similarly, a third destruction pulse is transmitted along the beamdirection, focused at the deepest focal point F₃ in the deepest focalzone Z₃. The pulse energy readily reaches the third focal zone due tothe earlier depletion of microbubbles at shallower depths. The thirddestruction pulse destroys the remaining microbubbles between the secondand third zones as indicated by line 292 in FIG. 14c, destroysmicrobubbles around the focal point as indicated by line 290, anddestroys a decreasing amount of microbubbles beyond the focal point F₃as indicated by line 294. A third interrogation pulse follows fordifferential detection of the microbubble destruction events in andaround zone Z₃.

In practice it has been found that peak microbubble destruction is notcentered exactly about the focal point axially, but in a depth regionjust prior to the focal point. This factor should be taken intoconsideration when considering the placement and overlap of multizonemicrobubble destruction regions.

The detected destruction events over the three zones are then combinedin accordance with the expression

|P_(F1)−P′_(F1)|+|P_(F2)−P′_(F2)|+|P_(F3)−P′_(F3)|

where P_(Fn) represents echoes following a destructive pulsetransmission to a given focal zone and P′_(Fn) represents the echoesfrom a subsequent interrogation pulse. The echoes from each focal zoneare spliced together to form a complete image line to the maximum depthof the image. In a preferred embodiment, instead of just detectingmicrobubble destruction events over the given focal zone, the techniqueconventionally used in multizone focus imaging, echoes are detected overthe full depth following each pulse. This enables the recording ofmicrobubble destruction events outside the given focal zone, providingthe greatest detection of destruction events. Thus, each pulse echo paircontains a line of echoes over the full image depth, which are thencombined to record the maximum number of microbubble destruction eventsfor the full image line.

It is also seen that, instead of transmitting a pair of pulses tointerrogate each focal zone, the echoes returned from later focal zonetransmissions can be combined with earlier echoes to differentiallydetect destruction events. That is, the first term of the aboveexpression could be |P_(F1)−P_(F2)|, for instance. However, the use ofpulse pairs for each focal zone is preferred, as the aperture changesaccompanying focal zone changes can deleteriously affect the precisionof the technique.

More uniform, artifact-free multizone microbubble destruction images canbe obtained by pulsing nonadjacent beams with time successive pulses.This ensures that each line of microbubbles will be approximatelyuniformly undisturbed at the beginning of the multizone sequence,preventing successions of bright and dim lines in the ultrasonic image.

FIGS. 15a, 15 b and 16 illustrate a preferred technique for displayingcontrast agent enhanced images when tissue perfusion is being observed.FIG. 16 illustrates a cross sectional view of the heart, including themyocardium 260 and the blood pool 250 within a chamber of the heart.When a contrast agent has been introduced into the bloodstream, a greatquantity of the agent will be contained within large blood pools such asthe heart chambers and major vessels, while only a relatively smallquantity of contrast agent will enter tissue and organs by way ofcapillary structures. In the heart image of FIG. 16, a large quantity ofcontrast agent will be present in the blood pool 250 while a lesseramount will be infused by capillary flow into the myocardium 260.

A conventional ultrasonic display of the cross sectional image of FIG.16 will cause pixels of greater signal level to be illuminated withgreater brightness or color. A typical display mapping characteristicwhich provides this result is shown in FIG. 15a by mappingcharacteristic 252. As detected pixel values increase, the displaypixels are shown with increasing brightness or color until reaching amaximum plateau level. As a result, the blood pool area 250 in FIG. 16will be shown brightly or highly colored, whereas the myocardium 260will be only dimly illuminated or colored.

When the myocardium is the area of interest in FIG. 16, a displaymapping characteristic such as that shown in FIG. 15b is employed. Thecurve 254 in this drawing is seen to begin at a zero level to suppressnoise in the image, then rises to a high level 256. Thereafter itdeclines to a level 258 for higher detected signal values. As a result,lower detected pixel values will be mapped to brightly illuminated orcolored display pixels, and higher detected pixel values will be mappedto more dimly illuminated or colored display pixel values. As aconsequence of this mapping, the myocardium 260 in FIG. 16 will bebrightly illuminated or colored, while the central blood pool is onlydimly colored or illuminated. This emphasis provides highlighting ofcontrast agent perfused tissue over blood pool areas.

Pulse transmission techniques can afford further improvement in contrastagent destruction and detection. While the exact physical mechanismscaused the by interaction of microbubbles with acoustic energy are quitecomplex, the sizes of microbubbles have an effect upon their destructionat certain frequencies. Since a microbubble contrast agent is oftencomprised of microbubbles of a wide range of diameters, microbubbledestruction events can be increased by transmitting a chirp ormultifrequency pulse. By transmitting a frequency modulated pulse, theprobability of transmitting destructive energy for a greater range ofmicrobubble sizes is increased. In addition, by modulating both thefrequency and amplitude of the destructive pulse, both microbubbledestruction and controlled oscillation can be induced. The initial highamplitude, low frequency period of the pulse, followed by a loweramplitude, higher frequency period can induce microbubble shelldestruction followed by oscillation of the released microbubble.

Another transmission technique which affords high pulse rates (PRF) isillustrated in FIGS. 17a-17 c. FIG. 17a illustrates the transmission ofa first pulse P₁ for contrast agent imaging of the heart, followed by asecond pulse P₂. In this example the pulses are transmitted at a lowPRF, and a significant period of time exists between the transmissiontimes of the pulses. During this time echoes 300 are first received fromcontrast agent in the myocardium, and later echoes 302 are received fromthe more distant pericardium. Differentiation of the echoes followingthe two pulses will detect the presence of contrast agent in themyocardium, followed by detection of the pericardium itself.

For procedures where it is only desirable to perform contrast agentimaging of the myocardium, a higher PRF transmission can be employed asshown in FIG. 17b. The higher PRF pulses have the unfortunate result ofartifact development. Echoes 300 return from the contrast agent in themyocardium following pulse P₁. But echoes 302 returning from thepericardium in response to the first pulse P₁ appear in the intervalfollowing the second pulse P₂ and can manifest themselves as an artifactin the image when echoes following the two pulses are differentiated. Toeliminate the artifact from the later returning echoes, incoherentdetection is employed prior to differentiation by the apparatus of FIG.1. As shown in FIG. 17c, incoherent detection and differentiationresults in positive polarity echoes 300′ from the myocardiummicrobubbles, and negative polarity echoes 302′ from the pericardium.The unwanted negative polarity echoes 302′ from the pericardium can thenbe removed by thresholding or clipping at the baseline, leaving only thedesired detection of the contrast agent in the myocardium.

What is claimed is:
 1. A method for ultrasonically imaging a region ofthe body which has been infused with a microbubble ultrasonic contrastagent comprising the steps of: transmitting a first pulse into the bodywhich is focused at a first focal point within the body to cause aresponse from microbubbles located over a first range of depths;receiving echoes following the transmission of said first pulse;transmitting a second pulse into the body which is focused at a secondfocal point within the body different from the first pulse to cause aresponse from microbubbles located over a range of depths correspondingat least partially to the first range; receiving echoes following thetransmission of said second pulse; and corresponding echoes received inresponse to the first and second pulses to produce an ultrasonic imageof the contrast agent.
 2. The method of claim 1, wherein saidtransmitting steps comprise transmitting first and second pulses alongsubstantially the same beam direction.
 3. The method of claim 2, whereinsaid transmitting steps cause microbubble response to comprise thedestruction of microbubbles.
 4. The method of claim 1, wherein timesuccessive pulses are transmitted in nonadjacent beam directions.
 5. Themethod of claim 4, wherein said microbubble response is a nonlinearresponse.
 6. The method of claim 1, wherein said microbubble response isa nonlinear response, and wherein said ultrasonic image is a harmonicultrasonic image.
 7. The method of claim 1, wherein said ultrasoniccontrast agent produces a harmonic response to ultrasonic pulsetransmissions.